Magnetic Resonance Imaging (MRI) can generate cross-sectional images in any plane (including oblique planes). Medical MRI most frequently relies on the relaxation properties of excited hydrogen nuclei (protons) in water and fat. When the object to be imaged is placed in a powerful, uniform magnetic field the spins of the atomic nuclei with non-integer spin numbers within the tissue all align either parallel to the magnetic field or anti-parallel. The output result of an MRI scan is an MRI contrast image or a series of MRI contrast images.
In order to understand MRI contrast, it is important to have some understanding of the time constants involved in relaxation processes that establish equilibrium following RF excitation. As the excited protons relax and realign, they emit energy at rates which are recorded to provide information about their environment. The realignment of proton spins with the magnetic field is termed longitudinal relaxation and the rate (typically about 1 s−1) required for a certain percentage of the tissue nuclei to realign is termed “R1 relaxation rate” or R1. T2-weighted imaging relies upon local dephasing of spins following the application of the transverse energy pulse; the transverse relaxation rate (typically >10 s−1 for tissue) is termed “R2 relaxation rate” or R2. These relaxation rates are also expressed as relaxation times T1 (=1/R1) and T2 (=1/R2). The total signal depends on the number of protons, or proton density PD. The total signal is decreased due to random motion of the protons, a process which can be enhanced by the application of a large bipolar gradient; moving protons acquire a phase difference which leads to a further signal loss. The signal loss indicates the diffusion of water molecules and can be measured as the apparent diffusion coefficient ADC. The measurement of the direction in which diffusion occurs is named fractional anisotropy FA. On the scanner console all available parameters, such as echo time TE, repetition time TR, flip angle α and the application of preparation pulses and gradients (and many more), are set to certain values. Each specific set of parameters generates a particular signal intensity in the resulting images depending on the characteristics of the measured tissue.
Generally MR images are qualitative in nature: the absolute image signal intensity has no meaning, it is the signal intensity differences, the contrast, which is interpreted. This leads to subjective reading of images, inherent inaccuracy and user dependence. MR quantification, on the other hand, aims at measurement of physical properties on an absolute scale. This provides a firm basis for objective measures and automated tissue recognition. Examples are measurement of brain volume for dementia follow-up, tumour volume for oncology and lesion load for Multiple Sclerosis.
There is a constant demand for improvements in MR imaging. It would therefore be desirable to provide improved and faster methods for obtaining measurements of physical properties such as R1, R2, PD and ADC.